I Field of the Invention
The present invention relates generally to providing biostable elastomeric coatings on the surfaces of implants which incorporate biologically active species having controlled release characteristics in the coating and relates particularly to providing a non-thrombogenic surface during and after timed release of the biologically active species. The invention is particularly described in terms of coatings on therapeutic expandable stent prostheses for implantation in body lumens, e.g., vascular implantation.
II Related Art
In surgical or other related invasive procedures, the insertion and expansion of stent devices in blood vessels, urinary tracts or other locations difficult to otherwise access for the purpose of preventing restenosis, providing vessel or lumen wall support or reinforcement and for other therapeutic or restorative functions has become a common form of long-term treatment. Typically, such prostheses are applied to a location of interest utilizing a vascular catheter, or similar transluminal device, to carry the stent to the location of interest where it is thereafter released to expand or be expanded in situ. These devices are generally designed as permanent implants which may become incorporated in the vascular or other tissue which they contact at implantation.
One type of self-expanding stent has a flexible tubular body formed of several individual flexible thread elements each of which extends in a helix configuration with the centerline of the body serving as a common axis. The elements are wound in the same direction but are displaced axially relative to each other and meet, under crossing a like number of elements also so axially displaced, but having the opposite direction of winding. This configuration provides a resilient braided tubular structure which assumes stable dimensions upon relaxation. Axial tension produces elongation and corresponding diameter contraction that allows the stent to be mounted on a catheter device and conveyed through the vascular system as a narrow elongated device. Once tension is relaxed in situ, the device at least substantially reverts to its original shape. Prostheses of the class including a braided flexible tubular body are illustrated and described in U.S. Pat. Nos. 4,655,771 and 4,954,126 to Wallsten and U.S. Pat. No. 5,061,275 to Wallsten et al.
Implanted stents have been used to carry medicinal agents, such as thrombolytic agents. U.S. Pat. No. 5,163,952 to Froix discloses a thermalmemoried expanding plastic stent device formulated to carry a medicinal agent in the material of the stent itself. Pinchuk, in U.S. Pat. No. 5,092,877, discloses a stent of a polymeric material which may have a coating associated with the delivery of drugs. Other patents which are directed to devices of the class utilizing bio-degradable or bio-sorbable polymers include Tang et al., U.S. Pat. No. 4,916,193, and MacGregor, U.S. Pat. No. 4,994,071.
A patent to Sahatjian, U.S. Pat. No. 5,304,121, discloses a coating applied to a stent consisting of a hydrogel polymer and a preselected drug such as a cell growth inhibitors or heparin. A further method of making a coated intravascular stent carrying a therapeutic material is described in Berg et al., U.S. Pat. No. 5,464,650, issued on Nov. 7, 1995 and corresponding to European Patent Application No. 0 623 354 A1 published Nov. 09, 1994. In that disclosure, a polymer coating material is dissolved in a solvent and the therapeutic material dispersed in the solvent; the solvent evaporated after application.
An article by Michael N. Helmus (a co-inventor of the resent invention) entitled xe2x80x9cMedical Device Designxe2x80x94A Systems Approach: Central Venous Cathetersxe2x80x9d, 22nd International Society for the Advancement of Material and Process Engineering Technical Conference (1990) relates to polymer/drug/membrane systems for releasing heparin. Those polymer/drug/membrane systems require two distinct types of layers to function.
It has been recognized that contacting blood with the surface of a foreign body in vivo has a tendency to induce thrombogenic responses and that as the surface area of a foreign device in contact with host blood increases, the tendency for coagulation and clot forming at these surfaces also increases. This has led to the use of immobilized systemic anti-coagulant or thrombolytic agents such as heparin on blood contacting surfaces such as oxygen uptake devices to reduce this phenomenon. Such an approach is described by Winters, et al., in U.S. Pat. Nos. 5,182,317; 5,262,451 and 5,338,770 in which the amine functional groups of the active material are covalently bonded using polyethylene oxide (PEO) on a siloxane surface.
Another approach is described in U.S. Pat. No. 4,613,665 to Larm in which heparin is chemically covalently bound to plastic surface materials containing primary amino groups to impart a non-thrombogenic surface to the material. Other approaches for bonding heparin are described in Barbucci, et al., xe2x80x9cCoating of commercially available materials with a new heparinizable materialxe2x80x9d, Journal of Biomedical Materials Research, Vol. 25, pp. 1259-1274 (1991); Hubbell, J. A., xe2x80x9cPharmacologic Modification of Materialsxe2x80x9d, Cardiovascular Pathology, Vol. 2, No. 3 (Suppl.), 121S-127S (1993); Gravlee, G. P., xe2x80x9cHeparin-Coated Cardiopulmonary Bypass Circuitsxe2x80x9d, Journal of Cardiothoracic and Vascular Anesthesia, Vol. 8, No. 2, pp. 213-222 (1994).
Moreover, drug elution rates for a coating containing a hydrophilic or a lipophobic drug is usually very fast initially when the coated device contacts body fluid or blood. One of the methods to reduce the so called xe2x80x9cburst effectxe2x80x9d is to add a membrane containing porosigen over the coating layer containing the biologically active material. See e.g., U.S. Pat. No. 5,605,696 to Eury et al. and U.S. Pat. No. to Helmus et al. U.S. Pat. No. 5,447,724. When the porosigen elutes, a porous membrane is formed and the drug in the undercoat will release. Even though the method might be quite successful to control the drug release, it increases the coating thickness, reduces the effective drug loading and introduces undesirable additional foreign materials into the patient. Hence, there is a need for a coating which reduces the burst effect but is not too thick and does not require the release of porosigens into the body.
With regard to stents, polymeric stents, although effective, may have mechanical properties that are inferior to those of metal stents of like thickness and weave. Metallic vascular stents braided of even relatively fine metal can provide a large amount of strength to resist inwardly directed circumferential pressure. A polymer material of comparable strength requires a much thicker-walled structure or heavier, denser filament weave, which in turn, reduces the cross-sectional area available for flow through the stent and/or reduces the relative amount of open space in the weave. Also, it is usually more difficult to load and deliver polymeric stents using catheter delivery systems.
While certain types of stents such as braided metal stents may be preferred for some applications, the coating and coating modification process of the present invention is not so limited and can be used on a wide variety of prosthetic devices. Thus, in the case of stents, the present invention also applies, for example, to the class of stents that are not self-expanding, including those which can be expanded, for instance, with a balloon; as well as polymeric stents of all kinds. Other medical devices that can benefit from the present invention include blood exchanging devices, vascular access ports, central venus catheters, cardiovascular catheters, extracorpeal circuits, vascular grafts, pumps, heart valves, and cardiovascular sutures, to name a few. Regardless of detailed embodiments, applicability of the invention should not be considered limited with respect to implant design, implant location or materials of construction. Further, the present invention may be used with other types of implantable prostheses.
Accordingly, it is a primary object of the present invention to provide a coating and process for coating a stent to be used as a deployed stent prostheses, the coating being capable of effective controlled long-term delivery of biologically active materials.
Another object of the invention is to provide a coating and process for coating a stent prostheses using a biostable hydrophobic elastomer in which biologically active species are incorporated within a coating.
Still another object of the present invention is to provide a multi-layer coating and process for the delivery of biologically active species in which the percentage of active material can vary from layer to layer.
Yet another object of the present invention is to provide a multi-layer coating and process for the delivery of biologically active species from a coating with a non-thrombogenic surface.
Another object of the invention is to provide a coating for the delivery of biologically active species having a top layer or topcoat which reduces the initial release of the species, in which the topcoat is substantially free of pores or porosigens and covers less than the entire surface of the undercoat. The topcoat can cover less than the entire surface of the undercoat before and/or while the device is implanted.
A further object of the invention is to provide a multilayer coating for the delivery of biologically active species such as heparin having a fluorosilicone top layer.
A still further object of the invention is to provide a multi-layer coating for the delivery of biologically active species such as heparin having a surface containing immobilized polyethylene glycol (PEG).
Other objects and advantages of the present invention will become apparent to those skilled in the art upon familiarization with the specification and appended claims.
The present invention provides a relatively thin layered coating of biostable elastomeric material containing an amount of biologically active material dispersed therein in combination with a non-thrombogenic surface that is useful for coating the surfaces of prostheses such as deployable stents.
The preferred stent to be coated is a self-expanding, open-ended tubular stent prostheses. Although other materials, including polymer materials, can be used, in the preferred embodiment, the tubular body is formed of a self expanding open braid of fine single or polyfilament metal wire which flexes without collapsing, readily axially deforms to an elongate shape for transluminal insertion via a vascular catheter and resiliently expands toward predetermined stable dimensions upon removal in situ.
In the process, the initial coating is preferably applied as a mixture, solution or suspension of polymeric material and finely divided biologically active species dispersed in an organic vehicle or a solution or partial solution of such species in a solvent or vehicle for the polymer and/or biologically active species. For the purpose of this application, the term xe2x80x9cfinally dividedxe2x80x9d means any type or size of included material from dissolved molecules through suspensions colloids and particulate mixtures. The biologically active material is dispersed in a carrier material which may be the polymer, a solvent, or both. The coating is preferably applied as a plurality of relatively thin layers sequentially applied in relatively rapid sequence and is preferably applied with the stent in a radially expanded state.
In many applications the layered coating is referred to or characterized as including an undercoat and topcoat. The coating thickness ratio of the topcoat to undercoat may vary with the desired effect and/or the elution system. Typically these are of different formulations with most or all of the biologically active material being contained in the undercoat and a non-thrombogenic or biocompatible non-porous surface found in the topcoat.
It is desirable that the topcoat be substantially free of connected pores or porosigens (materials which can elute during implantation and form pores). The addition of a porous membrane as a top coat will increase the coating thickness and reduce the overall drug loading. Also, the release of porosigens into the body can be undesirable since it introduces additional foreign materials into the body, which can cause the patient to have adverse reactions.
Since in some embodiments the topcoat should be substantially free of pores, the topcoat should cover less than the entire surface of the undercoat. Preferably, the topcoat should cover only about 10% to about 95% of the surface of the undercoat.
By partially covering the surface during manufacture or inducing xe2x80x9cbreaksxe2x80x9d in the topcoat during mounting/implanting of the coated device, the biologically active material or drug of the undercoat is permitted to be released from the undercoat through those parts of the undercoat which are not covered by the topcoat.
This mechanism is illustrated by FIG. 9, which shows a surface of a prosthesis 101 covered by a coating 102 comprising an undercoat 103 and a topcoat 104. The topcoat 104, which only partially covers the undercoat 103, leaves certain areas 106 of the undercoat, including drug particles 105, exposed to body fluids at the implantation site. It is through these xe2x80x9cuncoveredxe2x80x9d areas 106 of the undercoat that the drug particles 105 of the undercoat 103 are allowed to be released into the implantation site.
Additionally, it is preferred that the topcoats have an average thickness equivalent to the average particle size of the drug in the undercoat. Preferably the average thickness is about 1 to 7 microns and more preferable that the topcoat average thickness be about 1 to 5 microns. Also, the polymer of the topcoat may be the same as or different from the polymer of the undercoat.
The coating may be applied by dipping or spraying using evaporative solvent materials of relatively high vapor pressure to produce the desired viscosity and quickly establish coating layer thicknesses. The preferred process is predicated on reciprocally spray coating a rotating radially expanded stent employing an air brush device. The coating process enables the material to adherently conform to and cover the entire surface of the filaments of the open structure of the stent but in a manner such that the open lattice nature of the structure of the braid or other pattern is preserved in the coated device.
The coating is exposed to room temperature ventilation for a predetermined time (possibly one hour or more) for solvent vehicle evaporation. In the case of certain undercoat materials, thereafter the polymer material is cured at room temperature or elevated temperatures. Curing is defined as the process of converting the elastomeric or polymeric material into the finished or useful state by the application of heat and/or chemical agents which induce physico-chemical changes. Where, for example, polyurethane thermoplastic elastomers are used as an undercoat material, solvent evaporation can occur at room temperature rendering the undercoat useful for controlled drug release without further curing.
The applicable ventilation time and temperature for cure are determined by the particular polymer involved and particular drugs used. For example, silicone or polysiloxane materials (such as polydimethylsiloxane) have been used successfully. Urethane prepolymers can also be utilized. Unlike the polyurethane thermoplastic elastomers, some of these materials are applied as prepolymers in the coating composition and must thereafter be heat cured. The preferred silicone species have a relatively low cure temperatures and are known as a room temperature vulcanizable (RTV) materials. Some polydimethylsiloxane materials can be cured, for example, by exposure to air at about 90xc2x0 C. for a period of time such as 16 hours. A curing step may be implemented both after application of the undercoat or a certain number of lower layers and the top layers or a single curing step used after coating is completed.
The coated stents may thereafter be subjected to a postcure process which includes an inert gas plasma treatment, and sterilization which may include gamma radiation, ETO treatment, electron beam or steam treatment.
In the plasma treatment, unconstrained coated stents are placed in a reactor chamber and the system is purged with nitrogen and a vacuum applied to 20-50 mTorr. Thereafter, inert gas (argon, helium or mixture of them) is admitted to the reaction chamber for the plasma treatment. One method uses argon (Ar) gas, operating at a power range from 200 to 400 watts, a flow rate of 150-650 standard ml per minute, which is equivalent to about 100-450 mTorr, and an exposure time from 30 seconds to about 5 minutes. The stents can be removed immediately after the plasma treatment or remain in the argon atmosphere for an additional period of time, typically five minutes.
In accordance with the invention, the topcoat or surface coating may be applied in any of several ways to further control thrombolytic effects and optionally, control the release profile especially the initial very high release rate associated with the elution of heparin.
In one embodiment, an outer layer of fluorosilicone (FSi) is applied to the undercoat as a topcoat. The outer layer can also contain heparin. In another embodiment, polyethylene glycol (PEG) is immobilized on the surface of the coating. In this process, the underlayer is subjected to inert gas plasma treatment and immediately thereafter is treated by ammonia (NH3) plasma to aminate the surface. Amination, as used in this application, means creating mostly imino groups and other nitro containing species on the surface. This is followed by immediate immersion into electrophilically activated polyethylene glycol(PEG) solution with a reductive agent, i.e., sodium cyanoborohydride.
To form a topcoat which is substantially free of pores, porosigens or materials capable of eluting from the topcoat during implantation, should not be included in the composition used to form the topcoat. For example, a substantially non-porous topcoat can be produced from a topcoat composition which comprises a substantially pure polymeric material. The material preferably provides biocompatibility to the implanted device during and after release of the biologically active material.
To prepare a topcoat which covers less than the entire surface of the undercoat, a number of methods can be used. For instance, by controlling the thickness of the topcoat so that it has an average thickness less than that of the diameter of certain drug particles, the undercoat will be only partly covered by the topcoat since some of drug particles will not be covered by the topcoat.
Also, a partial topcoat can be formed by using a topcoat polymer which is incompatible with the undercoat polymer to generate a microphase separation in the topcoat. Furthermore, to make a topcoat which covers less than the entire surface of the undercoat or which only partially covers the undercoat, a poor solvent wash can be applied to the topcoat, to force the topcoat polymer to shrink so that the undercoat is not entirely covered.
In other embodiments the topcoat can partially or fully cover the undercoat prior to delivery or implantation of the device. The topcoat materials can be selected so they have certain water permeability. When water penetrates the topcoat and into the drug particles of the undercoat, the water will swell the particles or dissolve the particles. In either situation, it creates osmotic pressure in the surrounding coating material of the undercoat. The pressure then breaks the thinnest part of the topcoat, and leave the void space in the topcoat for the drug to elute out.
In another embodiment, the topcoat material has a different Young""s modulus (either while it is wet or dry) than that of the undercoat material. More specifically, the Young""s modulus can be higher for the topcoat material. During the mounting of the coated devices onto the delivery device or during deployment of the coated device, the coating undergoes compression or stretching. Topcoat materials with higher Young""s modulus tend to crack and form void spaces for the drug to elute from undercoat.
Another way to form a topcoat is to cover the undercoat with a bioabsorbable material. In this embodiment, the topcoat can cover either the entire undercoat or only part of the undercoat before or after implantation. Upon contact with body fluids, the bioabsorbable material of the topcoat will degrade. The rate of degradation depends upon the bioabsorbable material used. When the topcoat is partially absorbed, the undercoat is exposed to the body fluid and the drug is released, however the burst effect will be reduced.
The coated and cured stents having the modified outer layer or surface optionally are subjected to a final gamma radiation sterilization nominally at 2.5-3.5 Mrad. Argon (Ar) plasma treated stents enjoy full resiliency after radiation whether exposed in a constrained or non-constrained status, while constrained stents subjected to gamma sterilization without Ar plasma pretreatment lose resiliency and do not recover at a sufficient or appropriate rate where the undercoat is silicone.
The elastomeric materials that form the stent coating underlayers should possess certain properties. Preferably the layers should be of suitable hydrophobic biostable elastomeric materials which do not degrade. Surface layer or topcoat materials should minimize tissue rejection and tissue inflammation and permit encapsulation by tissue adjacent the stent implantation site. Exposed material is designed to reduce clotting tendencies in blood contacted and the surface is preferably modified accordingly. Thus, underlayers of the above materials are preferably provided with a silicone or a fluorosilicone outer coating layer which should not contain imbedded bioactive material, such as heparin in order to avoid the formation of pores in the topcoat. Alternatively, the outer coating may consist essentially of polyethylene glycol (PEG), polysaccharides, phospholipids, or combinations of the foregoing.
Polymers generally suitable for the undercoats or underlayers include silicones (e.g., polysiloxanes and substituted polysiloxanes), polyurethanes, thermoplastic elastomers in general, ethylene vinyl acetate copolymers, polyolefin elastomers, polyamide elastomers, and EPDM rubbers. The above-referenced materials are considered hydrophobic with respect to the contemplated environment of the invention. Surface layer or topcoat materials can include the same polymer as that of the undercoat. Examples of suitable polymers include without limitation fluorosilicones and polyethylene glycol (PEG), polysaccharides, phospholipids, and combinations of the foregoing.
While heparin is preferred as the incorporated active material, agents possibly suitable for incorporation include antithrobotics, anticoagulants, antibiotics antiplatelet agents, thrombolytics, antiproliferatives, steroidal and nonsteroidal antiinflammatories, agents that inhibit hyperplasia and in particular restenosis, smooth muscle cell inhibitors, growth factors, growth factor inhibitors, cell adhesion inhibitors, cell adhesion promoters and drugs that may enhance the formation of healthy neointimal tissue, including endothelial cell regeneration. The positive action may come from inhibiting particular cells (e.g., smooth muscle cells) or tissue formation (e.g., fibromuscular tissue) while encouraging different cell migration (e.g., endothelium) and tissue formation (neointimal tissue).
Suitable materials for fabricating the braided stent include stainless steel, tantalum, titanium alloys including nitinol (a nickel titanium, thermomemoried alloy material), and certain cobalt alloys including cobalt-chromium-nickel alloys such as Elgiloy(copyright) and Phynox(copyright). Further details concerning the fabrication and details of other aspects of the stents themselves, may be gleaned from the above referenced U.S. Pat. Nos. 4,655,771 and 4,954,126 to Wallsten and U.S. Pat. No. 5,061,275 to Wallsten et al., which are incorporated by reference herein.
Various combinations of polymer coating materials can be coordinated with biologically active species of interest to produce desired effects when coated on stents to be implanted in accordance with the invention. Loadings of therapeutic materials may vary. The mechanism of incorporation of the biologically active species into the surface coating, and egress mechanism depend both on the nature of the surface coating polymer and the material to be incorporated. The mechanism of release also depends on the mode of incorporation. The material may elute via interparticle paths or be administered via transport or diffusion through the encapsulating material itself.
For the purposes of this specification, xe2x80x9celutionxe2x80x9d is defined as any process of release that involves extraction or release by direct contact of the material with bodily fluids through the interparticle paths connected with the exterior of the coating. xe2x80x9cTransportxe2x80x9d or xe2x80x9cdiffusionxe2x80x9d are defined to include a mechanism of release in which the material released traverses through another material.
The desired release rate profile can be tailored by varying the coating thickness, the radial distribution (layer to layer) of bioactive materials, the mixing method, the amount of bioactive material, the combination of different matrix polymer materials at different layers, and the crosslink density of the polymeric material. The crosslink density is related to the amount of crosslinking which takes place and also the relative tightness of the matrix created by the particular crosslinking agent used. This, during the curing process, determines the amount of crosslinking and so the crosslink density of the polymer material. For bioactive materials released from the crosslinked matrix, such as heparin, a denser crosslink structure will result in a longer release time and reduced burst effect.
It will also be appreciated that an unmedicated silicone thin top layer provides some advantage and additional control over drug elution.